Magnetic resonance imaging (“MRI”) is a type of medical imaging modality for creating images of the inside of a human or animal body without using ionizing radiation such as x-rays. Mill uses a powerful magnet, such as a superconducting magnet, to create a strong, generally uniform, static magnetic field, usually referred to as “B0” or the main magnetic field, oriented along the longitudinal axis of the MRI scanner, conventionally referred to as the “z axis”. When a human body, or part of a human body, is placed in this main magnetic field, the nuclear spins that are associated with nuclei, such as hydrogen nuclei in tissue water (and to a lesser extent, other protons of other tissue types, such as lipids) become polarized. As a result, the magnetic moments that are associated with these spins become preferentially aligned along the direction of the main magnetic field, resulting in a small net tissue magnetization along the z axis. An MRI system also comprises components called gradient coils that produce smaller amplitude, spatially varying magnetic fields.
Typically, gradient coils are designed to produce magnetic field components that vary linearly in amplitude with position along one of the orthogonal x, y or z axes and, by convention, are referred to as Gx, Gy, and Gz magnetic field gradients. The Gx, Gy, and Gz magnetic field gradients from the gradient coils create various ramps on the corresponding magnetic field strengths, and concomitantly on the resonance frequencies of the nuclear spins, along these orthogonal axes, thereby spatially encoding the magnetic resonance (“MR”) signal by creating selected resonant frequencies or signal phases at various locations in the subject. By convention, the Gz magnetic field gradient is a slice-selection gradient, the Gy magnetic field gradient is phase-encoding gradient, and the Gx magnetic field gradient is a frequency encoding or read gradient. Radio frequency (“RF”) coils are used to create pulses of RF energy, as RF pulses interacting with the nuclear spins, at or near these various resonant frequencies. Subsequently, as the nuclear spins return to their equilibrium states (e.g., through T2 and T1 relaxation processes), an RF signal is emitted and detected by the MRI system.
In typical MR imaging, this process is repeated many times, at different phase-encoding gradients, for each selected slice or for multiple selected slices. Each of the corresponding received RF signals, with the influence from various magnetic field gradients, is then sampled, resulting in a corresponding line of “k-space”, also referred to herein as the “reception” k-space. Acquisition of multiple k-space lines is generally required to produce an image. As a result, image acquisition is comparatively slow, and is consequently subject to various artifacts, such as artifacts from motion of the patient within the MRI scanner. From reception k-space, which has spatial frequency information, and using a transformation (e.g., an inverse Fourier transformation) and any of various image reconstruction algorithms, a computing system converts the k-space information into a corresponding image for viewing.
Echo planar imaging (“EPI”) is a widely-used technique in MRI, especially to reduce image acquisition time, including multiple-shot EPI or single-shot EPI (“ss-EPI”). Single-shot echo planar pulse sequences are widely used in neuro-functional imaging, diffusion imaging, and perfusion imaging because of their fast acquisition speed, ability to freeze motion, and low specific absorption rate (“SAR”). These sequences, however, are highly susceptible to generating geometric distortions of the resulting image arising from, for example, static off-resonance effects such as magnetic susceptibility variations and B0-field inhomogeneity, or dynamic perturbations, including eddy currents. This distortion can be particularly severe in the frontal and temporal lobes of the brain, where the air-filled paranasal sinus and petrous portion of the temporal bone impose difficulties in applications such as functional MRI (“fMRI”) and diffusion imaging. Image distortion (Δd) in ss-EPI is given by (Equation 1):
      Δ    ⁢                  ⁢    d    =            γ              2        ⁢        π              ⁢          esp              Δ        ⁢                                  ⁢                  k          y                      ⁢    Δ    ⁢                  ⁢          B      0      where γ is the gyromagnetic ratio, esp is inter-echo spacing, ΔB0 is the change (i.e., inhomogeneity) of the main magnetic field B0, and Δky is the sampling interval in the phase-encoding direction of k-space (i.e., the distance between two consecutive k-space lines in the phase-encoding direction).
An obstinate and ongoing problem of EPI is image distortion, which is one of the primary obstacles preventing EPI from realizing its full potential as an imaging technology. As a consequence of this image distortion, EPI typically has been limited to low spatial resolutions. For example, to reduce distortion, echo spacing (esp) has been shortened by reducing the readout sampling points (kx), which unfortunately results in a lower spatial resolution. Other methods have been used to maintain or increase the spatial resolution, such as multi-shot EPI (e.g., Short Axis PROPELLER (SAP)-EPI and Readout-Segmented (RS)-EPI). These technologies, however, have other problems such as phase inconsistency among different k-space segments and/or requirements for correcting motion using sophisticated reconstruction algorithms. Another approach to reducing image distortion has involved increasing Δky, as demonstrated in parallel imaging methods such as SENSE and GRAPPA. The performance of distortion reduction in parallel imaging increases with the acceleration factor R, but at an expense of a decreased signal-to-noise ratio (SNR). At a larger R, however, gain in distortion reduction is offset by more SNR loss, which comes from the intrinsic SNR loss due to the reduced data sampling, the additional SNR loss arising from the parallel imaging reconstruction algorithm, and the imperfections of receiver coil geometry, which is described as a “g-factor” (g-factor ≧1). More specifically, the SNR is decreased by a factor of g-factor·√R. Because the g-factor is also a non-linear function of R and increases drastically for a higher acceleration factor, the acceleration factor R is typically limited to 2-3 for SNR considerations, limiting the amount of distortion reduction in parallel imaging.
Accordingly, there is an ongoing need for new technologies for reducing image distortion in EPI while achieving high spatial resolution. Such a technology should also maintain the fast acquisition speed of traditional EPI and, in addition, should be able to provide various imaging advantages, such as simultaneous suppression of lipid signals, and also use in anatomical regions that traditionally have not benefited from accurate imaging using conventional EPI techniques.